Magnetic resonance imaging (MRI) is a technique used frequently in medical settings to produce images of the inside of the human body. MRI is based on detecting nuclear magnetic resonance (NMR) signals, which are electromagnetic waves emitted by atomic nuclei in response to state changes resulting from applied electromagnetic fields. In particular, magnetic resonance (MR) techniques involve detecting NMR signals produced upon the re-alignment or relaxation of the nuclear spin of atoms in the tissue of the human body. MR techniques may be used to image and study the properties of tissue in a variety of regions of the human body, for example, for detection and/or diagnosis of tissue anomalies, study of blood flow, etc.
During an MRI procedure, NMR signals emitted from a volume of interest or from a slice (i.e., a relatively thin region) of the volume of interest are detected. The detected NMR signals may then be reconstructed to form a two-dimensional (2D) image of the slice. A 2D image is comprised of pixels, each pixel having an intensity (e.g., a magnitude or value) that is proportional to the strength of the NMR signal emitted by a corresponding location in the volume of interest. A plurality of such 2D images reconstructed from NMR signal data obtained from successive slices may be stacked together to form a three-dimensional (3D) image. A 3D image is composed of voxels, each voxel having an intensity proportional to the strength of the NMR signal emitted from a corresponding portion of the volume of interest.
As discussed above, MRI exploits the NMR phenomenon to distinguish various tissue characteristics. In particular, MRI operates by manipulating the spin characteristics of tissue, and more specifically, spin characteristics of hydrogen atoms (e.g., in water molecules) which compose a significant proportion of the human body, including both blood, tissue and fat. MRI techniques include aligning the spins of hydrogen nuclei with an axial magnetic field B0, and then perturbing the magnetic field in a targeted region with one or more radio frequency (RF) magnetic fields B1.
The NMR phenomenon results from exciting hydrogen nuclei by generating RF signals B1 at the Larmor frequency and applying them to a region of interest. The Larmor frequency is related to the rate at which nuclear spins precess about an axis at which the spins are aligned, which rate is, in turn, proportional to the strength of the axial magnetic field B0. When applied, the RF magnetic field B1 causes the nuclear spins to change orientation, and causes some nuclei to achieve a higher energy state.
When the RF signal B1 subsides, the nuclear spins realign with the axial magnetic field B0. Those nuclei that achieved the higher energy state upon excitation, return to the lower energy state by releasing electromagnetic energy. The released electromagnetic energy may be detected as NMR signals and used to form one or more images representative of the tissue type in the region of interest. The NMR signals may be detected using one or more RF coils sensitive to electromagnetic changes caused by the NMR signals, as discussed in further detail below.
The NMR phenomenon may be invoked in a number of ways. Conventional methods include applying the magnetic field B1 in a pulse sequence, referred to herein as an RF pulse sequence. One pulse sequence, commonly referred to as a spin-echo pulse sequence, includes applying an RF excitation signal (also known as an excitation pulse), followed by one or more RF refocusing signals (also known as refocusing pulses) to create a spin-echo, as described in connection with FIGS. 1A and 1B, which illustrates the magnetic field geometry in a coordinate frame 100.
With reference to the illustrated coordinate frame 100, the magnetic field B0 may be applied in substantial alignment with the z-axis. As will be appreciated by those skilled in the art, applying the magnetic field B0 along the z-axis will cause spins to align in one of two configurations: 1) a low energy configuration in which spins align in the −Z direction; and 2) a high energy configuration in which spins align in the +Z direction. The spin property of atomic nuclei can be viewed as a magnetic moment. When aligned in the +Z direction, the magnetic moment vector is similarly aligned in the +Z direction. Likewise, spins aligned in the −Z direction have a magnetic moment vector aligned in the −Z direction.
When a region of interest is placed in magnetic field B0, more spins will align in the +Z direction than in the −Z direction (i.e., more spins will orient themselves according to the high energy configuration). As a result, the region of interest will have a net magnetization vector M oriented in the Z direction. That is, the net magnetization vector M will have a relatively large Mz component, and substantially zero Mx and My components. The magnitude of the Mz component will depend, at least in part, on the strength of magnetic field B0. The configuration illustrated in FIG. 1A is referred to as the equilibrium state.
When an appropriate RF pulse is applied at 90° to the z-axis (e.g., applied substantially parallel to the XY plane, as illustrated by magnetic field B1 in FIG. 1B), spins may be perturbed from their equilibrium state in alignment with the axial magnetic field B0 such that the net magnetization vector M is rotated into the XY plane. FIG. 1B illustrates the magnetic field geometry after a 90° excitation pulse has been applied, and the spins have been rotated 90° from their alignment with the axial magnetic field B0 (Z-axis) to be aligned with the Y-axis. As shown, the component Mz goes substantially to zero, and the Mxy component (referred to as the transverse component) becomes relatively large.
With the net magnetization vector aligned in the XY plane (referred to as the transverse plane), the spins will precess about the Z-axis at the Larmor frequency. That is, the net magnetization vector will rotate about the Z-axis as illustrated by arrow 121 in FIG. 1B. After the RF pulse subsides and is no longer influencing the region of interest, the spins begin to re-align with magnetic field B0, returning to the equilibrium state along the z-axis. The time constant governing how the net magnetization vector returns to equilibrium is referred to as the spin relaxation time (T1). That is, T1 describes the rate at which the net magnetization vector regains its equilibrium component in the direction of magnetic field (e.g., the rate at which the Mz component of the net magnetization is recovered).
As discussed above, the application of the 90° RF pulse causes the net magnetization vector to rotate to the XY plane. As a result, the net magnetization vector obtains a substantial component Mxy. As the net magnetization vector precesses about the z-axis, the spins begin to dephase at different locations within the excited region. There are a number of reasons for the dephasing. First, atomic nuclei are influenced by slightly different magnetic field strengths (e.g., as a result of inhomogeneities in the applied magnetic fields), resulting in the various spins precessing at slightly different Larmor frequencies. In addition, the chemical environment may cause spins to precess at different rates, and magnetic field gradients used to localize NMR effects may also contribute to spin dephasing. As a result of dephasing, the component of the net magnetization vector in the XY plane (i.e., Mxy) decays to zero or substantially zero. The time constant governing the decay of Mxy is referred to as the transverse relaxation or spin-spin relaxation time (T2). T2 depends at least upon various molecular interactions and generally microscopic inhomogeneities in the applied magnetic fields.
Following the 90° excitation pulse, a 180° RF pulse (referred to as a refocusing pulse) may be applied to cause the phases to regain coherence, recovering the transverse magnetization (e.g., recovering Mxy). The recovery of the transverse magnetization subsequently produces an NMR signal referred to as a spin-echo. The spin-echo may be detected to characterize the subject matter of the region of interest (e.g., to ascertain the hydrogen concentration characteristic of particular tissues in the body). It should be appreciated that spin lattice and spin-spin relaxation occurs simultaneously. That is, spin dephasing occurs simultaneously with realignment of the net magnetization vector with the z-axis, though at different rates. In particular, T2 can be no larger and is typically substantially smaller than T1.
Spin-echo pulse sequences typically include applying a 90° RF pulse followed by one or more 180° RF pulses as illustrated in FIGS. 2A and 2B. In particular, one conventional spin-echo pulse sequence is illustrated in FIG. 2A, wherein a 90° RF excitation pulse is configured to rotate the net magnetization vector M into the XY plane, where the spins begin to dephase causing the transverse magnetization Mxy to decay. After some interval, a 180° RF refocusing pulse may be applied to align the spin phases. As result of the refocusing pulse, a spin-echo S is generated which can be detected by RF coils (e.g., the RF coils configured to deliver the RF pulse sequences). At some later point related to T1, the magnetization vector recovers the equilibrium state and the process (e.g., an excitation pulse, followed by a refocusing pulse) may be repeated to obtain sufficient NMR data to characterize a region being imaged.
One variant on the spin-echo pulse sequence is referred to as the fast spin-echo (FSE) sequence. FSE exploits the fact that T2 is typically substantially smaller than T1. Accordingly, a plurality of refocusing pulses may be applied between successive applications of excitation pulses, as illustrated in FIG. 2B. In particular, the FSE pulse sequence may include an initial RF excitation signal (e.g., a 90° pulse) followed by a series of RF refocusing signals (e.g., a series of 180° pulse). As a result, multiple spin-echoes may be detected for each excitation pulse, increasing the speed of image acquisition.